In general, the healing of wounds, burns, and other injuries is an uncertain endeavor. The clinician cannot be certain about the condition of the tissue being treated, the efficacy of treatments, and whether further treatments or a change in treatments is appropriate. As a particular example, many chronic wounds, such as pressure ulcers or venous stasis ulcers often linger for months or even years, despite the various treatments being applied. These wounds are particularly intractable for a variety of reasons; with age, nutrition, diabetes, infection, marginalized immune systems, and other factors, contributing to the ongoing difficulties in healing. In most cases, such wounds are chronic because the wound healing is stalled relative to one or more aspects of the process. In such circumstances, it is not unusual for the clinician to be unsure about the status of the wounded tissue, at what point in wound healing the tissue is held up, and what new treatment modality should be applied.
In general, the detection and diagnosis of deep tissue injuries (chronic wounds, bruises, surgical wound complications, sports injuries, etc.) is difficult. There are a variety of methods and devices that can be utilized to aid ongoing diagnosis. For example, tissue biopsies can be taken, and used in tissue cultures and histology. These traditional methods are disadvantaged by the time delay in evaluating tissue cultures or histology, which easily can be a week or two. Additionally, these approaches are invasive, and actually cause further damage to the tissue. As histology relies on thin slices of tissue, which are dyed and examined optically with a microscope, histology typically only provides a direct indication of the tissue structure in two-dimensions.
Alternate technologies have been developed for non-invasive histology or tissue imaging, including X-ray, magnetic resonance imaging (MRI and NMR), computed axial tomography (CAT) scanning, and positron emission tomography (PET). These technologies are used for a variety of applications (mammography, brain scans, etc.), but are seldom used for examining soft tissue wounds, and can expose the patient to high-energy radiation (x-rays, etc.). Ultrasound, which is widely used for pre-natal examination, can also be used for examining wounds. In particular, Longport Inc. (Glen Mills, Pa.) offers a high frequency (20 MHz) ultrasound scanner, described in U.S. Pat. No. 6,073,045 (Dyson et al.) that scans tissue to modest depths (2 cm) but with “high” spatial resolution (65 microns).
However, because many biological structures, including cells, are much smaller than 65 microns, there is a need for other imaging technologies that offer higher resolution, but at a lower cost than MRI and some other medical imaging technologies. There are a variety of technologies, including confocal scanning microscopy, optical coherence tomography (OCT), diffuse optical tomography (DOT), second harmonic generation (SHG) and multi-photon based microscopy, which apply optical techniques to obtain high resolution images either in-vivo or in-vitro. In these cases, imaging resolution can be a few microns or less, which certainly enables detection of much finer structures than does ultrasound. On the other hand, light absorption and scatter limit the imaging depth in most tissues to only ˜1-4 mm. While many of these systems have been used for optical histology, providing a two-dimensional image of the tissue, newer technologies, such as full field OCT systems, enable three-dimensional images. Although these various optical imaging technologies are being used for both research and clinical diagnosis, in general there are opportunities for improvements to facilitate wider diagnostic use, including the application in wound assessment.
It appears that the ideal diagnostic device for diagnostic assessment of deep tissue injuries does not yet exist. Such a device would provide sharp and clear images with an imaging resolution of 2-5 μm within a relatively deep imaging depth of 1-5 cm, while covering a sizeable imaging area of 1-100 cm2, with a cost of $10,000-$50,000, depending on features. There are some emerging technologies, such as terahertz imaging or the “NMR Mouse,” which may yet fulfill this need. However, lacking such a definitive technical solution, then it appears that dual modality devices offer the greatest potential. For example, a dual ultrasound and optical imaging system may work, if these imaging modalities could be synergistically and inexpensively combined. Alternately, there appears to be opportunity in multi-modal optical devices, which combine various optical imaging modalities (such as photographic imaging, confocal microscopy, OCT, or DOT) in one device. As these various technologies all manipulate light to create images, judicious design choices could facilitate useful combinations. Even so, optical imaging in scattering tissues is then necessarily limited to imaging depths in the 1-6 mm range. Nonetheless, imaging within that depth range could still have considerable value.
Obviously the device requirements derive from the physiological and optical properties of the tissues in question. In the particular case of wound assessment, it is necessary to understand both the physiobiology and the optical properties of wounds. As wounds heal, they normally progress through a sequence of overlapping interactive phases, starting with coagulation and progressing through inflammation, proliferation (which includes granulation, angiogenesis, and epithelialization), and remodeling.
Success in wound healing is very much dependent on the rebuilding of the extra-cellular matrix (ECM), which is initially dependent on fibroblasts. Fibroblasts migrate into the wound site, and begin to build the ECM by depositing a protein called fibronectin. Fibronectin is deposited with some directionality, mirroring the axis of the fibroblasts. The fibroblasts then produce collagen, with the collagen deposition generally aligned to the fibronectin pattern. Over time, fibronectin is replaced by Type III collagen and ultimately by Type I collagen. In parallel, angiogenesis occurs, and new capillaries bud and grow into the collagen network, creating granulation tissue. Over time, granulation tissue continues to change, attempting to become as much like normal tissue as possible. For example, as the wound contracts, and is subsequently remodeled and influenced by stresses from neighboring tissues, the collagen becomes increasingly organized. Even late in the remodeling phase, which can end six months to a year post injury, collagen in a scar will be replaced and rearranged as the wound attempts to regain its original function.
In considering the in-vivo optical imaging of wounds, the fact that both collagen and capillaries are optically birefringent, represents an opportunity to monitor the wound healing processes involving the formation and remodeling of the extra-cellular matrix (ECM) and granulation tissue. Obviously, a medical optical imaging system could have polarization sensitivity to help see these features. Additionally however, granulation tissues and wound tissues have other attributes, such as altered optical transmission properties and cellular and extra-cellular morphologies which could effect optical imaging therein, which a properly design medical imaging system could utilize, if it were designed correctly.
As mentioned previously, various optical technologies (confocal microscopy, OCT, SHG, fluoroscopy, diffuse optical tomography) and combinations thereof have been developed for use in medical tissue imaging. Confocal microscopy and OCT are particularly of interest, as these two technologies have been specifically developed to enable optical imaging “deep” into tissue, which is a turbid media in which scattering severely limits the potential for tissue imaging. Both confocal microscopes and OCT systems are often designed to image sub-cellularly, so that internal cellular structures such as the nuclei and mitochondria can be examined. To provide the desired submicron (˜0.2-1.0 μm) resolution, these systems utilize very fast optics (Numerical Aperture (NA) ˜0.8-1.4), often enabled by immersion optics. As a result, both the field of view and imaging depth of such systems are constrained, thereby limiting the in-vivo imaging utility of these devices.
Presently, OCT systems are used more widely than are confocal imaging systems, because they can image to greater depths (˜2-3×) into tissues. However, as tissue scattering limits OCT imaging, the technology has been most successfully applied in ocular applications, where the tissue is weakly scattering, to examine visual pathologies such as glaucoma, diabetic retinopathy, macular degeneration, etc. In particular, the coherence/interference effect utilized by OCT provides greater signal discrimination (rejection of out of focus light) than does confocal microscopy, which relies on one or more pinholes for signal discrimination. In general, an OCT system is basically a fiber optic based interferometer, typically using a low coherence (broad band, for example ˜30-70 nm) light source. Such systems are provided with a sampling arm, which includes a fiber optic probe to direct light onto the tissue. These system also have a reference fiber optic arm with a retro-reflector. The interference effect between the sample arm light and the reference arm light allows OCT systems to control the depth of focus, so that a small longitudinal distance is in focus. Images are constructed by first measuring the in-depth profile of the backscattered light intensity in the axial (depth) direction. This backscattered light is predominately that from a single scattering event, with a lesser contribution from light that encountered small angle scattering events. In-depth profiling is performed by measuring the echo time delay and intensity of backscattered or reflected light. Distance or spatial information is determined from the time delay of reflected echoes. To create a two-dimensional image, the fiber optic beam is moved laterally across the surface (x-axis) and in-depth profiles (z-axis) are obtained at discrete points along the surface. The net result is that the resolution (1-20 microns) and dynamic range of the sample are in-focus and enhanced as compared to the portion of the sample the un-focused beam traveled through. This can be particularly advantageous for imaging in turbid, light scattering optical media, such as tissue. However, OCT imaging depth and resolution, and signal strength are all effected by the scattering properties of the tissue being examined. In general, the less scattering there is (smaller scattering coefficient, μs), the deeper the imaging. However, the directionality of the scattering (forward or back) also effects signal strength, signal localization, and resolution. Exemplary OCT system patents include U.S. Pat. Nos. 5,659,392 and 6,034,774 (both to Marcus et al.), both of which are assigned to the same assignee as the present invention.
Polarization sensitive OCT systems have also been developed. An exemplary prior art system, described in U.S. Pat. No. 6,208,415 (DeBoer et al.), has been used at Massachusetts General Hospital to examine dermal tissues, burns, scars, and tendons. Another exemplary prior art OCT system, described in U.S. Pat. No. 6,615,072 (Izatt et al.) is equipped with a polarization compensation system, so as to desensitize the device to polarization degradation effects that occur in bent single mode optical fibers. Another similar system is a polarization sensitive low coherence reflectometer, such as described in U.S. Pat. No. 5,459,570 (Swanson et al.) which has 11 micron resolution and 120 dB signal to noise ratio. Although the fiber optic OCT systems can have a small probe for in-vivo testing, these systems are complicated and expensive, and are not likely to be used by a clinician in wound assessment either in the field or in many clinical settings.
OCT systems can also be designed with more traditional optics (rather than fiber optics), by combining the attributes of a Michelson interferometer with those of a microscope. In particular, wide field (or full field) of view OCT systems have been developed, wherein a microscope objective lens is used to illuminate and image the sample, while a second lens is placed in the reference arm of the interferometer. An exemplary wide field OCT system is described in U.S. Pat. No. 6,940,602 (Dubois). While OCT systems are advantaged over confocal systems relative to imaging depth and depth resolution, the systems are more complicated (with the reference arm) and potentially less flexible.
In confocal microscopy, light is directed through a pinhole to create a spot of light, which is projected or imaged into the sample under examination. Returning, backscattering image light is imaged to a pinhole stop located in an intermediate image plane. The image light is directed to a sensor, to provide data signals. As a result, only light from the focal plane can reach the detector. Other potential image locations within the sample, such as out-of-focus planes or spatially offset locations within the same plane are blocked out, as the spatial filtering effect of the pinhole acts as an intensity-gate. This results in an “optical section.” With the confocal microscope, the z-resolution, or optical sectioning thickness, depends on a number of factors, including wavelength λ, pinhole size, numerical aperture (NA) of the objective lens, refractive index (n) of the components, and the alignment of the instrument. Whereas, viewing depth largely depends on tissue scattering and absorption properties and pinhole size.
The confocal microscope, as described in U.S. Pat. No. 3,013,467, was originally developed by Marvin Minsky as an approach to examine each point of a specimen and measure the amount of light scattered or absorbed by that point, while minimizing the collection and detection of light scattered by neighboring points. Most simply, confocal microscopes are constructed with a single pinhole that defines both the size of the illumination light that will be imaged to the tissue, and the size of the spot of return light allowed to reach the detector. Alternately, the optical design can provide separate pinholes in the illumination and imaging paths. However, the price of single-point illumination is being able to measure only one point at a time. Thus, in a traditional confocal microscope, the specimen is scanned point by point and the resulting image is reconstructed thereafter. The sample can be moved relative to the microscope either by laterally translating the sample itself (with translation stages) or sweeping (with galvanometers) the illuminating light beam over the sample.
However, the utility of confocal microscopes can be limited by weak signals. Compared to a normal microscope, the amount of light that is seen in the final image is greatly reduced by the pinhole, sometimes up to 90-95%. To compensate for this loss of light somewhat, lasers are used as light sources instead of the conventional mercury arc lamps because they produce extremely bright light at very specific wavelengths. As an example, U.S. Pat. No. 5,032,720 (White) describes a beam scanning confocal fluorescence microscope in which the light emitted from an argon laser is focused to the sample, and the coherence of the laser allows it to act as its own pinhole, so that an illuminating pinhole is not needed. As a result, the optical system has greater optical efficiency, and stronger signals will be available at the detector.
As one approach to increase the throughput of a confocal microscope, both the stage and the light source can be kept stationary, while the specimen is scanned with an array of light spots transmitted through apertures. Alternately, a time variant array of light spots is created by spinning a Nipkow disk within the microscope assembly. For example, U.S. Pat. No. 4,802,748 (McCarthy et al.) describes a tandem scanning reflected light confocal microscope in which a Nipkow disc has a series of apertures located in an annular pattern of spiral arms on the disc surface. U.S. Pat. No. 5,067,805 (Corle) describes a polarization sensitive confocal scanning optical microscope with a spinning Nipkow disc in which the polarization beamsplitter is tilted relative to the optical system to prevent crosstalk from stray reflected light.
As an alternate approach to improving the throughput of a confocal microscope, systems have been developed wherein the mechanically rotating Nipkow disc has been replaced by a spatial light modulator array, which can be electrically addressed and thus function as a programmable pinhole array. In particular, U.S. Pat. No. 5,587,832 (Krause) describes a spatially light modulated confocal microscope in which a modulator array, such as a liquid crystal device (LCD), a digital micro-mirror device (DMD), or a micro-shutter array is image conjugate to the tissue and is operated to function as a programmable multi-pinhole generator. In the system of Krause '832, two modulator arrays, one for illumination, and a second for detection, are used in tandem, under the control of a central processor. While Krause '832 provides the basic elements of a modulator based confocal imaging system, the various approaches described therein lack optical design attributes (such as telecentricity, uniform flood illumination, and focus adjustments) that would improve the performance and utility of the concept. Additionally, the various designs lack the polarization sensitive optics, and control thereof, that would be useful in examining extra-cellular structures (such as collagen) in normal tissues (such as skin), wounded tissues, and granulation (healing) tissues. Krause '832 also does not consider a device with multiple imaging modalities that could be enabling for examining deep tissue injury.
In a second patent, U.S. Pat. No. 5,923,466 (Krause), another version of a spatially light modulated confocal microscope system is described, in which a single DMD array is used in dual roles as both the source pinhole generator and the detector pinhole generator. Within a large reflective optical system, used both in collection and detection. The system of Krause '466 is fairly complicated, using a dual Offner type reflective imaging optics and a single mode optical fiber couple laser source, and is not designed either for low cost and ease of use, nor for examining wounds and optically birefringent tissue structures.
U.S. Pat. No. 5,867,251 (Webb) describes a tandem scanning confocal ophthalmoscope utilizing two spatial light modulators to create an image of an object plane located within the interior of the eye. This system is similar to that of Krause '832, but is optimized for ocular diagnostic applications rather than for looking at skin, wounds, and birefringent tissue structures, and thus lacks many of the same attributes discussed with respect to Krause '832 above.
By comparison, U.S. Pat. No. 6,399,935 (Jovin et al.) attempts to improve upon Krause '832 by providing an alternate DMD-based programmable confocal microscope with improved light efficiency and dual confocal and non-confocal (conventional) microscopy capability. In particular, this patent uses the on-state pixels to collect the confocal image and the off-state pixels to collect the non-confocal image, using either two detector arrays or two light sources to provide the duality of use. Additionally, pseudorandom pixel patterns, for example based upon cyclic Hadamard matrices, are suggested as a means to improve capture speed and the effective light efficiency. U.S. Pat. No. 6,483,641 (MacAulay) describes further DMD-based programmable microscopes, but ones in which the modulator arrays are located in optical planes conjugate to the aperture stop of the system, rather than conjugate to the object and image planes of the system. In this instance, the intent is to provide rapid control of the angular spectrum of the illumination light that is incident to the sample. A further reference, U.S. Pat. No. 6,144,489 (Wilson), describes a confocal microscope in which an encoded mask is used in the illumination system with patterns to generate combined confocal and non-confocal images. The mask can be a spinning disc, or a spatial light modulator, such as a DMD or a ferroelectric liquid crystal device. Again, all of the above prior art devices lack the appropriate design attributes for a medical imaging system that is optimized for use in examining skin, wounds, deep tissue injury, and birefringent tissue structures.
Wilson has also reported the use of structured illumination to provide a wide field-of-view confocal-like optical sectioning capability, without using either a Nipkow disc or a modulator array to address the specimen. Alternately, some systems have been proposed, such as those described in U.S. Pat. Nos. 6,769,769 and 6,927,860 (both to Podoleanu et al.), in which both OCT and confocal microscopy are combined together in a dual modality instrument, so that both types of images can be captured sequentially or simultaneously. However, these devices utilize single point imaging and scanning, rather than an array imaging or wide field of view approach.
It is noted that some portable, non-OCT or non-confocally based, optical devices for tissue diagnosis using polarization optics have also been developed. As an example, Lekam Medical (Devon, United Kingdom) offers the Cytoscan, which uses orthogonal polarization spectral imaging technology developed by Cytometrics Inc., and described in a U.S. Pat. No. 5,983,120 (Groner et al.); U.S. Pat. Nos. 6,438,396 and 6,650,916 (both to Cook et al.). This system is designed to provide images of the micro-circulatory vascular network, and is not optimized to examine the collagen network present in the dermal layers of skin. The Cytoscan system does not provide the proper optical wavelengths, high contrast polarizers, polarization control, or depth imaging to properly examine wounds and granulation tissues.
In considering the need for tissue imaging systems, which would be appropriate for wound assessment and other similar purposes, and which would be capable of imaging tissue structures with large fields of view at various depths, with the option of polarization sensitivity, it is seen that the range of present devices do not fulfill the anticipated diagnostic needs. In particular, there are needs for design for compact multi-functional diagnostic medical imaging systems. There are also opportunities for improved medical imaging devices that offer a wide range of capabilities and operational modalities.